Methods of fabricating stents with enhanced fracture toughness

ABSTRACT

Stents and methods of manufacturing a stents with enhanced fracture toughness are disclosed.

CROSS-REFERENCE

This is a continuation of application Ser. No. 12/845,536 filed on Jul.28, 2010 which is a continuation of application Ser. No. 12/772,698filed on May 3, 2010, now U.S. Pat. No. 8,323,329, which is a divisionalapplication of Ser. No. 11/454,968 filed on Jun. 15, 2006, now U.S. Pat.No. 7,731,890, all of which are incorporated by reference herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to methods of fabricating stents having selectedmechanical properties.

2. Description of the State of the Art

This invention relates to radially expandable endoprostheses, which areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel.

A stent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices, which function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of the diameter of a bodily passage ororifice. In such treatments, stents reinforce body vessels and preventrestenosis following angioplasty in the vascular system. “Restenosis”refers to the reoccurrence of stenosis in a blood vessel or heart valveafter it has been treated (as by balloon angioplasty, stenting, orvalvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through a bodily lumen to a region, such as alesion, in a vessel that requires treatment. “Deployment” corresponds tothe expanding of the stent within the lumen at the treatment region.Delivery and deployment of a stent are accomplished by positioning thestent about one end of a catheter, inserting the end of the catheterthrough the skin into a bodily lumen, advancing the catheter in thebodily lumen to a desired treatment location, expanding the stent at thetreatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon. The stent is thenexpanded by inflating the balloon. The balloon may then be deflated andthe catheter withdrawn. In the case of a self-expanding stent, the stentmay be secured to the catheter via a retractable sheath or a sock. Whenthe stent is in a desired bodily location, the sheath may be withdrawnwhich allows the stent to self-expand.

The stent must be able to satisfy a number of mechanical requirements.First, the stent must be capable of withstanding the structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel. Therefore, a stent must possess adequate radialstrength. Radial strength, which is the ability of a stent to resistradial compressive forces, is due to strength and rigidity around acircumferential direction of the stent. Radial strength and rigidity,therefore, may also be described as, hoop or circumferential strengthand rigidity.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.For example, a radially directed force may tend to cause a stent torecoil inward. Generally, it is desirable to minimize recoil.

In addition, the stent must possess sufficient flexibility to allow forcrimping, expansion, and cyclic loading. Longitudinal flexibility isimportant to allow the stent to be maneuvered through a tortuousvascular path and to enable it to conform to a deployment site that maynot be linear or may be subject to flexure. Finally, the stent must bebiocompatible so as not to trigger any adverse vascular responses.

The structure of a stent is typically composed of scaffolding thatincludes a pattern or network of interconnecting structural elementsoften referred to in the art as struts or bar arms. The scaffolding canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. The scaffolding is designed so that the stent can beradially compressed (to allow crimping) and radially expanded (to allowdeployment). A conventional stent is allowed to expand and contractthrough movement of individual structural elements of a pattern withrespect to each other.

Additionally, a medicated stent may be fabricated by coating the surfaceof either a metallic or polymeric scaffolding with a polymeric carrierthat includes an active or bioactive agent or drug. Polymericscaffolding may also serve as a carrier of an active agent or drug.

Furthermore, it may be desirable for a stent to be biodegradable. Inmany treatment applications, the presence of a stent in a body may benecessary for a limited period of time until its intended function of,for example, maintaining vascular patency and/or drug delivery isaccomplished. Therefore, stents fabricated from biodegradable,bioabsorbable, and/or bioerodable materials such as bioabsorbablepolymers should be configured to completely erode only after theclinical need for them has ended.

However, there are potential shortcomings in the use of polymers as amaterial for implantable medical devices, such as stents. There is aneed for a manufacturing process for a stent that addresses suchshortcomings so that a polymeric stent can meet the clinical andmechanical requirements of a stent.

SUMMARY OF THE INVENTION

Certain embodiments of the present invention include a stent comprisinga cylindrically aligned bending element formed by a first bar arm and asecond bar arm, the angle between the bar arms being greater than about90°, wherein the stent is fabricated from a tube radially expanded by atleast about 400%.

Further embodiments of the present invention include a stent comprisinga cylindrically aligned bending element formed by a first bar arm and asecond bar arm, an angle between each of the bar arms and thecircumferential direction being less than about 45°, wherein the stentis fabricated from a tube radially expanded by at least 500%.

Additional embodiments of the present invention include a stentcomprising a plurality of cylindrically aligned bending elements, theangles between the bending elements being greater than about 90°.

Other embodiments of the present invention include a method offabricating a stent comprising: radially expanding a tube to at leastabout 400%; and cutting a pattern comprising a cylindrically alignedbending element formed by a first bar arm and a second bar arm, theangle between the bar arms being greater than about 90°, wherein thestent is fabricated from a tube radially expanded by at least about400%.

Some embodiments of the present invention include a method forfabricating a stent comprising: conveying a gas into a poly(L-lactide)tube disposed within a cylindrical mold to increase a pressure insidethe tube, wherein the increased pressure radially expands the polymerictube to conform to the inside surface of the mold; applying tensionalong the axis of the tube to axially extend the tube; and fabricating astent from the radially expanded and axially extended tube.

Certain embodiment of the present invention include a method forfabricating a stent comprising: processing a polymer form to increasethe Tg of the polymer at least about 10° C.; and fabricating a stentfrom the processing form.

Additional embodiments of the present invention include a method forfabricating a stent comprising: processing a polymer form so as toincrease the Tg of the polymer to at least about 40° C. above ambienttemperature to allow storage of the processed polymer at the ambienttemperature; and fabricating a stent from the processed polymer.

Other embodiments of the present invention include a method forfabricating a stent comprising: processing a polymer form so as toincrease the Tg of the polymer to at least about 20° C. above a crimpingtemperature.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIGS. 2A-C depict blow-molding of a polymeric tube.

FIG. 3 depicts an exemplary stent pattern.

FIG. 4 depicts a bending element from the pattern in FIG. 3.

FIG. 5 depicts an alternative stent pattern.

FIGS. 6-8 depict images of expanded stents.

FIGS. 9-10 depict graphs of differential scanning calorimetry results.

DETAILED DESCRIPTION OF THE INVENTION

The various embodiments of the present invention relate to polymericstents and methods of fabricating polymeric stents with favorablemechanical properties. The present invention can be applied to devicesincluding, but is not limited to, self-expandable stents,balloon-expandable stents, stent-grafts, and grafts (e.g., aorticgrafts).

A stent can have a scaffolding or a substrate that includes a pattern ofa plurality of interconnecting structural elements or struts. FIG. 1depicts an example of a view of a stent 100. Stent 100 has a cylindricalshape with an axis 160 and includes a pattern with a number ofinterconnecting structural elements or struts 110. In general, a stentpattern is designed so that the stent can be radially compressed(crimped) and radially expanded (to allow deployment). The stressesinvolved during compression and expansion are generally distributedthroughout various structural elements of the stent pattern. The presentinvention is not limited to the stent pattern depicted in FIG. 1. Thevariation in stent patterns is virtually unlimited.

The underlying structure or substrate of a stent can be completely or atleast in part made from a biodegradable polymer or combination ofbiodegradable polymers, a biostable polymer or combination of biostablepolymers, or a combination of biodegradable and biostable polymers.Additionally, a polymer-based coating for a surface of a device can be abiodegradable polymer or combination of biodegradable polymers, abiostable polymer or combination of biostable polymers, or a combinationof biodegradable and biostable polymers.

A stent such as stent 100 may be fabricated from a polymeric tube or asheet by rolling and bonding the sheet to form a tube. A stent patternmay be formed on a polymeric tube by laser cutting a pattern on thetube. Representative examples of lasers that may be used include, butare not limited to, excimer, carbon dioxide, and YAG. In otherembodiments, chemical etching may be used to form a pattern on a tube.

The pattern of stent 100 in FIG. 1 varies throughout its structure toallow radial expansion and compression and longitudinal flexure. Apattern may include portions of struts that are straight or relativelystraight, an example being a portion 120. In addition, patterns mayinclude bending elements 130, 140, and 150.

Bending elements bend inward when a stent is crimped to allow radialcompression. Bending elements also bend outward when a stent is expandedto allow for radial expansion. After deployment, a stent is under staticand cyclic compressive loads from the vessel walls. Thus, bendingelements are subjected to deformation during use. “Use” includes, but isnot limited to, manufacturing, assembling (e.g., crimping stent on acatheter), delivery of stent into and through a bodily lumen to atreatment site, and deployment of stent at a treatment site, andtreatment after deployment.

As indicated above, a stent has certain mechanical requirements. A stentmust have sufficient radial strength to withstand structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel. In addition, the stent must possess sufficientflexibility to allow for crimping, expansion, and cyclic loading. Also,a sufficiently low profile, that includes diameter and size of struts,is important. As the profile of a stent decreases, the easier is itsdelivery, and the smaller the disruption of blood flow.

Polymers tend to have a number of shortcomings for use as materials forstents. One such shortcoming is that many biodegradable polymers have arelatively low modulus, and thus relatively low radial strength.Compared to metals, the strength to weight ratio of polymers is smallerthan that of metals. A polymeric stent with inadequete radial strengthcan result in mechanical failure or recoil inward after implantationinto a vessel. To compensate for the relatively low modulus, a polymericstent requires significantly thicker struts than a metallic stent, whichresults in an undesirably large profile.

Another shortcoming of polymers is that many polymers, such asbiodegradable polymers, tend to be brittle under physiologicalconditions or conditions within a human body. Specifically, suchpolymers can have a Tg, which is defined below, above human bodytemperature which is approximately 37° C. These polymer systems exhibita brittle fracture mechanism in which there is little or no plasticdeformation prior to failure. As a result, a stent fabricated from suchpolymers can have insufficient toughness for the range of use of astent. In particular, it is important for a stent to be resistant tofracture throughout the range of use of a stent, i.e., crimping,delivery, deployment, and during a desired treatment period.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. Tg of a given polymer can be dependent on the heating rate andcan be influenced by the thermal history of the polymer. Furthermore,the chemical structure of the polymer heavily influences the glasstransition by affecting mobility.

Other potential problems with polymeric stents include creep, stressrelaxation, and physical aging. Creep refers to the gradual deformationthat occurs in a polymeric construct subjected to an applied load. It isbelieved that the delayed response of polymer chains to stress duringdeformation causes creep behavior. Creep occurs even when the appliedload is constant. Creep can cause an expanded stent to retract radiallyinward, reducing the effectiveness of a stent in maintaining desiredvascular patency. The rate at which polymers creep depends not only onthe load, but also on temperature. In general, a loaded construct creepsfaster at higher temperatures.

Stress relaxation is also a consequence of delayed molecular motions asin creep. Contrary to creep, however, which is experienced when the loadis constant, stress relaxation occurs when deformation (or strain) isconstant and is manifested by a reduction in the force (stress) requiredto maintain a constant deformation

Physical aging, as used herein, refers to densification in the amorphousregions of a semi-crystalline polymer. Physical aging ofsemi-crystalline polymers that have glass transition temperatures (Tg)above their normal storage temperature, which, for the purposes of thisinvention is room temperature, i.e., from about 15° C. to about 35° C.,occurs primarily through the phenomenon known as densification.Densification occurs when polymer chains rearrange in order to move froma non-equilibrium state to an equilibrium state. The reordering ofpolymer chains tends to increase the modulus of the polymer resulting ina brittle or more brittle polymer.

Thus, physical aging results in an increase in brittleness of a polymerwhich can result in cracking of struts upon crimping and deployment.Since physical aging results from densification of amorphous regions ofa polymer, an increase in crystallinity can reduce or inhibit physicalaging.

However, it is well known by those skilled in the art that themechanical properties of a polymer can be modified through variousprocessing techniques, such as, by applying stress to a polymer. JamesL. White and Joseph E. Spruiell, Polymer and Engineering Science, 1981,Vol. 21, No. 13. The application of stress can induce molecularorientation along the direction of stress which can modify mechanicalproperties along the direction of applied stress. For example, strengthand modulus are some of the important properties that depend uponorientation of polymer chains in a polymer. Molecular orientation refersto the relative orientation of polymer chains along a longitudinal orcovalent axis of the polymer chains.

A polymer may be completely amorphous, partially crystalline, or almostcompletely crystalline. A partially crystalline polymer includescrystalline regions separated by amorphous regions. The crystallineregions do not necessarily have the same or similar orientation ofpolymer chains. However, a high degree of orientation of crystallitesmay be induced by applying stress to a semi-crystalline polymer. Thestress may also induce orientation in the amorphous regions. An orientedamorphous region also tends to have high strength and high modulus alongan axis of alignment of polymer chains. Additionally, for some polymersunder some conditions, induced alignment in an amorphous polymer may beaccompanied by crystallization of the amorphous polymer into an orderedstructure. This is known as stress induced crystallization.

As indicated above, due to the magnitude and directions of stressesimposed on a stent during use, it is important for the mechanicalstability of the stent to have suitable mechanical properties, such asstrength and modulus, in the axial and circumferential directions.Therefore, it can be advantageous to modify the mechanical properties ofa tube, to be used in the fabrication of a stent, by induced orientationfrom applied stress in the axial direction, circumferential direction,or both. Since highly oriented regions in polymers tend to be associatedwith higher strength and modulus, it may be desirable to incorporateprocesses that induce alignment of polymer chains along one or morepreferred axes or directions into fabrication of stents.

Therefore, it can be desirable to fabricate a stent from a polymerictube with induced orientation in the axial direction and in thecircumferential direction. A biaxial oriented tube may be configured tohave desired strength and modulus in both the circumferential and axialdirections.

The degree of radial expansion, and thus induced radial orientation andstrength, of a tube can be quantified by a radial expansion (RE) ratio:

$\frac{{Outside}\mspace{14mu} {Diameter}\mspace{14mu} ({OD})\mspace{14mu} {of}\mspace{14mu} {Expanded}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Inside}\mspace{14mu} {Diameter}\mspace{14mu} ({ID})\mspace{14mu} {of}\mspace{14mu} {Tube}}$

The RE ratio can also be expressed as a percent expansion:

% Radial expansion=(RE ratio−1)×100%

Similarly, the degree of axial extension, and thus induced axialorientation and strength, may be quantified by an axial extension (AE)ratio:

$\frac{{Length}\mspace{14mu} {of}\mspace{14mu} {Extended}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Length}\mspace{14mu} {of}\mspace{14mu} {Tube}}$

The AE ratio can also be expressed as a percent expansion:

% Axial expansion=(AE ratio−1)×100%

In some embodiments, a polymeric tube may be deformed by blow molding.In blow molding, a tube can be deformed or expanded radially byincreasing a pressure in the tube by conveying a fluid into the tube.The polymer tube may be deformed or extended axially by applying atensile force by a tension source at one end while holding the other endstationary. Alternatively, a tensile force may be applied at both endsof the tube. The tube may be axially extended before, during, and/orafter radial expansion.

In some embodiments, blow molding may include first positioning a tubein a cylindrical member or mold. The mold may act to control the degreeof radial deformation of the tube by limiting the deformation of theoutside diameter or surface of the tube to the inside diameter of themold. The inside diameter of the mold may correspond to a diameter lessthan or equal to a desired diameter of the polymer tube. Alternatively,the fluid temperature and pressure may be used to control the degree ofradial deformation by limiting deformation of the inside diameter of thetube as an alternative to or in combination with using the mold.

The temperature of the tube can be heated to temperatures above the Tgof the polymer during deformation to facilitate deformation. The polymertube may also be heated prior to, during, and subsequent to thedeformation. In one embodiment, the tube may be heated by conveying agas above ambient temperature on and/or into the tube. The gas may bethe same gas used to increase the pressure in the tube. In anotherembodiment, the tube may be heated by translating a heating element ornozzle adjacent to the tube. In other embodiments, the tube may beheated by the mold. The mold may be heated, for example, by heatingelements on, in, and/or adjacent to the mold.

Certain embodiments may include first sealing, blocking, or closing apolymer tube at a distal end. The end may be open in subsequentmanufacturing steps. The fluid, (conventionally a gas such as air,nitrogen, oxygen, argon, etc.) may then be conveyed into a proximal endof the polymer tube to increase the pressure in the tube. The pressureof the fluid in the tube may act to radially expand the tube.

Additionally, the pressure inside the tube, the tension along thecylindrical axis of the tube, and the temperature of the tube may bemaintained above ambient levels for a period of time to allow thepolymer tube to be heat set. Heat setting may include maintaining a tubeat a temperature greater than or equal to the Tg of the polymer and lessthan the Tm of the polymer for a selected period to time. The selectedperiod of time may be between about one minute and about two hours, ormore narrowly, between about two minutes and about ten minutes.

In heat setting, the polymer tube may then be cooled to below its Tgeither before or after decreasing the pressure and/or decreasingtension. Cooling the tube helps insure that the tube maintains theproper shape, size, and length following its formation. Upon cooling,the deformed tube retains the length and shape imposed by an innersurface of the mold.

Properties of a polymer such as fracture toughness are affected by theoverall degree of crystallinity and the number and size of crystaldomains in a semi-crystalline polymer. It has been observed thatfracture toughness is increased by having a large number of smallcrystal domains in a polymer surrounded by an amorphous domain. Such acrystal structure can also reduce or prevent creep, stress relaxation,and physical aging. In some embodiments, the size of crystal domains maybe less than 10 microns, 4 microns, or, more narrowly, less than 2microns. The overall crystallinity may be less than 50%, 40% or, morenarrowly, less than 20%.

In certain embodiments, the temperature of the deformation processand/or heat setting can be used to control the crystallinity to obtainthe desired crystal structure described above. In general,crystallization tends to occur in a polymer at temperatures between Tgand Tm of the polymer and it varies with temperature in this range. Insome embodiments, the temperature can be in a range in which the crystalnucleation rate is larger than the crystal growth rate. In oneembodiment, the temperature can be in a range in which the crystalnucleation rate is substantially larger than the crystal growth rate.For example, the temperature can be where the ratio of the crystalnucleation rate to crystal growth rate is 2, 5, 10, 50, 100, or greaterthan 100. In another embodiment, the temperature range may be in rangebetween about Tg to about 0.2(Tm−Tg)+Tg.

FIGS. 2A-C illustrate an embodiment of blow molding a polymer tube foruse in manufacturing a stent. FIG. 2A depicts an axial cross-section ofa polymer tube 200 with an outside diameter 205 positioned within a mold210. FIG. 2B depicts a radial cross-section of polymer tube 200 and mold210. Mold 210 may act to limit the radial deformation of polymer tube200 to a diameter 215, the inside diameter of mold 205. Polymer tube 200may be closed at a distal end 220. Distal end 220 may be open insubsequent manufacturing steps. A fluid may be conveyed, as indicated byan arrow 225, into an open proximal end 230 of polymer tube 200. Atensile force 235 is applied at proximal end 230 and a distal end 220.

Polymer tube 200 is heated by heating nozzles 260 on a support 265 thatblow a heated gas as shown by arrows 270. Support 265 translates backand forth along the axis of the mold as shown by arrows 275 and 280. Theincrease in pressure inside of polymer tube 200, facilitated by anincrease in temperature of the polymer tube, causes radial deformationof polymer tube 200, as indicated by an arrow 240. FIG. 2C depictspolymer tube 200 in a deformed state with an outside diameter 245 withinmold 210.

To illustrate the importance of orientation in a stent pattern, FIG. 3depicts an exemplary stent pattern 300 for use with embodiments of apolymeric tube or a sheet. In an embodiment, stent pattern 300 can becut from a polymeric tube using laser machining Stent pattern 300 isshown in a flattened condition so that the pattern can be clearlyviewed. When the flattened portion of stent pattern 300 is in acylindrical form, it forms a radially expandable stent.

As depicted in FIG. 3, stent pattern 300 includes a plurality ofcylindrical rings 305 with each ring including a plurality of diamondshaped cells 310. Embodiments of stent pattern 300 may have any numberof rings 305 depending on a desired length of a stent. For reference,line A-A represents the longitudinal axis of a stent using the patterndepicted in FIG. 3. Diamond shaped cells 310 include bending elements315 and 320. Stent pattern 300 can also includes bending elements 325and 330. The angles of bending elements 315, 320, 325, and 330correspond to angles θ₁, θ₂, θ₃, and θ₄. Angles θ₁, θ₂, θ₃, and θ₄ are42, 42, 41, and 21 degrees, respectively. Diamond shaped cells 310 aremade up of bar arms 335 and 340 that form bending element 315 and bararms 345 and 350 that form bending element 320.

When stent 300 is crimped, bending elements 315, 320, 325, and 330 flexinward and angles θ₁, θ₂, θ₃, and θ₄ decrease, allowing the stent to beradially compressed. With respect to bending elements 315, 320, and 325,struts on either side of the bending elements bend toward each other.However, in bending element 330, the strut of the diamond-shaped elementtends to bend toward the linking strut which tends to remain relativelyparallel to the longitudinal axis during crimping.

Pattern 300 further includes linking arms 355 that connect adjacentcylindrical rings. Linking arms 355 are parallel to line A-A and connectadjacent rings between intersection 360 of cylindrically adjacentdiamond-shaped elements 310 of one ring and intersection 360 ofcylindrically adjacent diamond shaped elements 310 of an adjacent ring.As shown, linking elements connect every other intersection along thecircumference.

The curved portions of bending elements experience substantial stressand strain when a stent is crimped and deployed. Therefore high strengthand toughness are very important in these regions. For example, aclose-up view of bending element 315 is depicted in FIG. 4 to illustratethe direction of stress in a bending element. Compressive and outwardradial stress on a stent cause substantially no strain in straightsections 400. However, such radial stresses result in relatively highstress and strain in curved portion 410 of bending element 315. Forexample, when a stent is expanded, angle θ₁ of bending element 315increases. The region above a neutral axis 415 experiences relativelyhigh compressive stress and strain and the region below neutral axis 415experiences relatively high tensile stress and strain. Alternatively,when a stent is crimped, angle θ₁ of bending element 315 decreases andthere is tensile stress and strain above neutral axis 415 andcompressive stress and strain below neutral axis 415.

The tensile and compressive strain follow the axis or curvature ofbending element 315, for example, line 420. Ideally, the most effectiveorientation to improve fracture toughness is along the length of theaxis of the strut. However, radial expansion imparts orientation andfracture toughness along the circumferential direction, as shown by lineB-B. An angle φ between a point on the axis of the stent and thecircumferential direction B-B tends to decrease moving along bendingelement 315 from the straight sections 400 to an apex 425 of bendingelement 315.

An exemplary stent having the pattern of FIG. 3 can be cut from apoly(L-lactide) (PLLA) tube that is about 0.084 in inside diameter. Adesired crimped diameter may be about 0.055 in and an expanded diameterabout of 0.134 in. Such a stent can be fabricated from an extruded tubethat is radially expanded between 200% and 400%. For a stent with thepattern shown in FIG. 3, and the dimensions provided above, cracks havebeen observed to form in the curved portion of bending elements uponexpansion of the stent to the expanded diameter.

For a given radius of curvature, increasing angle θ₁ of bending element315 tends to increase angle φ along the axis of bending element 315,making bending element 315 along curved portion 410 closer inorientation with the circumferential direction B-B. As a result, thestrength and toughness of bending element 315 are increased when thereis induced radial orientation in the stent. The relative orientation ofpoints along the axis, angle φ, of a bending element also depends on theradius of curvature. Increasing the radius of curvature of bendingelement 315 also makes bending element 315 along curved portion 410closer in orientation with the circumferential direction B-B.

Therefore, it is advantageous to decrease the relative orientationbetween the axis of bar arms or struts in curved portions and thecircumferential direction in a fabricated stent. Certain embodiments ofthe invention include stents having bending elements with angles greaterthan about 80°, or more narrowly, greater than about 90°, or 110°. Thestent may have an uncrimped or fabricated diameter that allows the stentto be crimped to a selected crimped diameter at which the bendingelements have an angle between 0° to 50°, or more narrowly between 0° to30°.

FIG. 5 depicts a stent pattern 500 similar to pattern 300 in FIG. 3. Theangles of bending elements 515, 520, 525, and 530 are about 113°, 113°,116°, and 55°, respectively. Therefore, the orientation of points on theaxis of the bending elements of pattern 500 are closer to thecircumferential direction than that in stent pattern 300. The radii ofcurvature of bending element 515 and 520 can be between about 0.014 inand 0.02 in. The radii of curvature of bending element 525 can bebetween about 0.009 in and 0.013 in. The radii of curvature of bendingelement 525 can be between about 0.0026 in and 0.0035 in.

In an embodiment, the outside diameter (OD) of a fabricated stent can bebetween 0.07 in and 0.165 in. The crimped diameter of a stent havingstent pattern 500 may be less than 0.06 in, 0.036 in, 0.032 in, or morenarrowly less than 0.028 in.

In certain embodiments, it may be advantages to fabricate a stent from atube that has been radial expanded to greater than 400%. As indicatedabove, cracks have been observed in high strain regions of stentfabricated from a tube expanded in the 200% to 400% range. In someembodiments, a stent may be fabricated from a tube that has been radialexpanded to greater than 500%, 600%, 700%, or greater than 800%. Thetube may be used to fabricate stents having a variety of patterns. Insome embodiments, a stent with a stent pattern 500 can be fabricatedfrom tube radially expanded to greater than 400%.

Such a stent may then show a greater increase in fracture toughness andstress over a stent fabricated from a tube radially expanded in a rangebetween 200% and 400%. As a result, such a stent may have fewer or nocracks when expanded to an intended deployment diameter. Increasing thedegree of expansion tends to impart greater strength and toughness.Thus, increasing the degree of expansion may extend the range of adiameter that a stent can be deployed.

Exemplary process conditions for expanding a PLLA tube between 400% and700% include a temperature of heated air at the heat nozzle between 205°F. and 285° F. The heat nozzle air flow rate can be between about 60 and65 SCFH (standard cubic feet per hour). The pressure of nitrogenconveyed into the tube can be between 177 psi and 250 psi. The tensionapplied axially to extend the tube can be between about 75 g and 105 g.

The advantages of expanding in a range greater than 400% is shown by thefollowing example. A PLLA tube was extruded to an ID of 0.024 in and anOD of 0.074 in. The extruded tubing was radially expanded using blowmolding 470% to an ID of 0.125 in and OD of 0.137 in. Five stents wereprepared from the expanded tubing. The expanded tubing was laser cut toform a stents with a pattern similar to stent pattern 500 in FIG. 5. Thestents were crimped, mounted on a catheter, and sterilized with E-beamradiation. The stents were expanded by a balloon on the catheter in a37° C. water bath to 0.138 in. The stents were removed and examined.FIGS. 6 and 7 show images of a stent expanded to 0.138 in. FIG. 6depicts the entire stent and FIG. 7 depicts a close-up view. The stentappears to be substantially free of cracks.

The stents were placed on another catheter and expanded further to 0.158in. FIG. 8 depicts an image of this expanded stent which shows cracksforming in the high strain regions. The images demonstrate theeffectiveness of increased biaxial orientation for the PLLA system.

As shown above, radial expansion above 400% can increases fracturetoughness of an expanded stent. Radial expansion above 400% can alsoaddress other issues with polymeric stents, such as stent retentionduring crimping and physical aging during long term storage.

As discussed above, physical aging results in an increase in brittlenessof a polymer which can result in cracking of struts upon crimping anddeployment. Polymeric stents generally are stored below ambienttemperatures to reduce or prevent physical aging the polymer that cancause cracking in stent struts during crimping and deployment. Stentscan be stored in freezers at temperatures below 0° C. Storing thepolymeric stents at low temperature reduces the segmental motions ofpolymer chains that result in densification.

In general, it would be desirable to store a polymeric stent close toambient temperature. However, many polymers have Tg's low enough toallow significant long term aging or densification to occur during thetime frame of long term storage, which can be a few days, a month, 3months, 6 months, or more than 6 months. Although Tg is defined as thetemperature at which the onset of segmental motion in the chains of thepolymer occurs, the glass transition is not sharp or discontinuous for apolymer with amorphous regions. Rather, there is a gradual transitionfrom the brittle to the ductile state corresponding to a gradualincrease in segmental motion. Thus, even for polymers with Tg's aboveambient temperatures, significant physical aging can occur during longterm storage. Increasing the difference between the storage temperatureand the Tg reduces the segmental motion of polymer chains which reducesor eliminate the effects of long term aging.

In addition, crimping of a polymeric stent at ambient temperatures canresult in an outward recoil of the stent from the crimped radius,reducing stent retention on the catheter. Due to shape memory of thepolymer, the stent recoils outward toward the fabricated diameter.

Such outward recoil can be reduced by heating the stent above ambienttemperatures during crimping. However, it has been observed thatelevated crimping temperatures can result in fracture of struts duringcrimping and upon deployment. Specifically, a PLLA stent fabricated froma polymeric tube expanded 300% from an extruded tube that is crimped at50° C. results in fracture during deployment. This observed increase inmechanical damage to the stent is a result of stress relaxation of thepolymer during the crimping process, due to the crimping being conductedclose to the Tg of the polymer. This stress relaxation will result ingreater experienced stress during the expansion of the stent duringdeployment. This will, in turn, result in a greater probability ofcracking during the expansion of the stent.

Increasing the difference between the elevated crimping temperature andthe Tg reduces the likelihood of cracking of struts.

In general, deforming a polymer form or construct can increase the Tg ofthe polymer. The increased order from orientation and inducedcrystallization caused by deformation tends to increase the temperaturenecessary for segmental motion of polymer chains, which corresponds toTg.

For a given polymer system, the degree of deformation, or specifically,expansion of a polymeric tube, may be correlated with an increase in Tg.Thus, an increase in Tg can allow storage of the polymer form at ahigher temperature with little or no negative effects of physical aging,or other visco-elastic phenomena. For example, the Tg can be increasedto allow storage at ambient temperature. In addition, the Tg can beincreased to allow crimping at a selected elevated temperature withoutcracking of stent struts.

In certain embodiments, a stent can be fabricated from a polymeric tubethat allows crimping at a selected elevated temperature with no orsubstantially no cracking of struts. The polymeric tube can be radiallyexpanded to a degree of expansion that allows crimping at the elevatedtemperature. The degree of expansion can be between 200% and 400%. Inother embodiments, the degree of expansion can be between 400% and 800%.The selected elevated temperature can be at least 10° C., 20° C., 30°C., 40° C., or 50° C. below the Tg of the polymer.

In additional embodiments, a stent can be fabricated from a polymerictube that allows long term storage at a selected temperature. Forexample, the temperature can be at or near an ambient temperature. Thepolymeric tube can be radially expanded to a degree of expansion thatallows storage at the selected temperature with little or no negativeeffects of physical aging. As above, the degree of expansion can bebetween 200% and 400%. In other embodiments, the degree of expansion canbe between 400% and 800%. The storage temperature can at least 30° C.,40° C., 50° C., 60° C., or 70° C. below the Tg of the polymer.

Differential scanning calorimetry (DSC) was used to study the increasein the Tg due the radial orientation induced by radial expansion. Ingeneral, DSC is a technique that may be used to identify thermaltransitions in a polymer. Thermal transitions include, for example,crystallization and melting. A thermal transition in a polymer may beendothermic (sample absorbs heat) or exothermic (sample expels heat).Glass and melting transitions are exothermic and crystallization isendothermic.

In a typical DSC run, a polymer sample is heated at a constant rate. Theheat inflow or outflow into the sample is controlled to keep the heatingrate constant. When the sample undergoes a thermal transition, heat iseither absorbed or expelled. At the glass transition and meltingtransition, heat flow into the sample decreases. When a polymer samplecrystallizes, the heat flow into the sample increases.

The Tg of PLLA tubes was studied at 300% and 500% radial expansion. DSCruns were performed for two samples for each degree of expansion. For500% radial expansion PLLA tubing was extruded to an ID of 0.021 in andan OD of 0.072 in. For 300% radial expansion, PLLA tubing was extrudedto an ID of 0.018 in and an OD of 0.056 in. The extruded tubing wasradially expanded using blow molding.

FIG. 9 depicts the results of DSC runs for samples expanded to 300%.Curve 900 corresponds to the first sample and curve 905 corresponds tothe second sample. Troughs 901 and 906 depict the glass transition,which is about 62° C. in each case. In addition, peaks 902 and 907correspond to the crystallization transition of the polymer for thefirst and second samples, respectively.

The melted samples at the end of each run were quenched to a solid form.DSC runs were then performed on the quenched samples for comparison.These samples correspond to PLLA without induced orientation. Curve 910corresponds to the first sample and curve 915 corresponds to the secondsample.

FIG. 10 depicts the results of DSC runs for samples expanded to 500%.Curve 1000 corresponds to the first sample and curve 1005 corresponds tothe second sample. Troughs 1001 and 1006 depict the glass transition,which is about 71° C. in each case. Curves 1000 and 1005 do not havepeaks analogous to peaks 902 and 907 in FIG. 9. This indicates thatpolymer of the samples expanded 500% was completely or almost completelycrystallized due to stress induced crystallization. The highcrystallinity reduces physical aging. The melted samples at the end ofeach run were quenched to a solid form. DSC runs were then performed onthe quenched samples for comparison. These samples correspond to PLLAwithout induced orientation. Curve 1010 corresponds to the first sampleand curve 1015 corresponds to the second sample.

Thus, the Tg increased from 62° C. to 71° C. from 300% to 500% radialexpansion. A stent fabricated from a tube expanded 500% was crimped at50° C. without strut fracture. Also, it is expected that the increase inTg allows for an increase in storage temperature.

Polymers can be biostable, bioabsorbable, biodegradable or bioerodable.Biostable refers to polymers that are not biodegradable. The termsbiodegradable, bioabsorbable, and bioerodable are used interchangeablyand refer to polymers that are capable of being completely degradedand/or eroded when exposed to bodily fluids such as blood and can begradually resorbed, absorbed, and/or eliminated by the body. Theprocesses of breaking down and eventual absorption and elimination ofthe polymer can be caused by, for example, hydrolysis, metabolicprocesses, bulk or surface erosion, and the like.

It is understood that after the process of degradation, erosion,absorption, and/or resorption has been completed, no part of the stentwill remain or in the case of coating applications on a biostablescaffolding, no polymer will remain on the device. In some embodiments,very negligible traces or residue may be left behind. For stents madefrom a biodegradable polymer, the stent is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished.

Representative examples of polymers that may be used to fabricate orcoat an implantable medical device include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate),poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(caprolactone), poly(trimethylene carbonate), polyester amide,poly(glycolic acid-co-trimethylene carbonate), co-poly(ether-esters)(e.g. PEO/PLA), polyphosphazenes, biomolecules (such as fibrin,fibrinogen, cellulose, starch, collagen and hyaluronic acid),polyurethanes, silicones, polyesters, polyolefins, polyisobutylene andethylene-alphaolefin copolymers, acrylic polymers and copolymers otherthan polyacrylates, vinyl halide polymers and copolymers (such aspolyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether),polyvinylidene halides (such as polyvinylidene chloride),polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such aspolystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose. Another type of polymer based on poly(lacticacid) that can be used includes graft copolymers, and block copolymers,such as AB block-copolymers (“diblock-copolymers”) or ABAblock-copolymers (“triblock-copolymers”), or mixtures thereof.

Additional representative examples of polymers that may be especiallywell suited for use in fabricating or coating an implantable medicaldevice include ethylene vinyl alcohol copolymer (commonly known by thegeneric name EVOH or by the trade name EVAL), poly(butyl methacrylate),poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508,available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidenefluoride (otherwise known as KYNAR, available from ATOFINA Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethyleneglycol.

The examples and experimental data set forth above are for illustrativepurposes only and are in no way meant to limit the invention. Thefollowing examples are given to aid in understanding the invention, butit is to be understood that the invention is not limited to theparticular materials or procedures of examples.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

1. A stent comprising a cylindrically aligned bending element formed bya first bar arm and a second bar arm, the angle between the bar armsbeing greater than about 90°, wherein the stent is fabricated from atube radially expanded by at least about 400%.
 2. The method of claim 1,wherein the radius of curvature of the bending element is between 0.014in and 0.02 in.
 3. The method of claim 1, wherein the angle between thebar arms is greater than about 100°.
 4. The method of claim 1, whereinthe angle between the bar arms is greater than about 110°.
 5. The methodof claim 1, wherein the polymer comprises a biostable polymer,biodegradable polymer, or a combination thereof.
 6. A stent comprising acylindrically aligned bending element formed by a first bar arm and asecond bar arm, an angle between each of the bar arms and thecircumferential direction being less than about 45°, wherein the stentis fabricated from a tube radially expanded by at least 500%.
 7. Thestent of claim 6, wherein the bending element comprises poly(L-lactide).8. The stent of claim 6, wherein the bending element comprises abiodegradable polymer, a biostable polymer, and/or a combination of botha biodegradable and biostable polymer.
 9. A stent comprising a pluralityof cylindrically aligned bending elements, the angles between thebending elements being greater than about 90°.
 10. A method offabricating a stent comprising: radially expanding a tube to at leastabout 400%; and cutting a pattern comprising a cylindrically alignedbending element formed by a first bar arm and a second bar arm, theangle between the bar arms being greater than about 90°, wherein thestent is fabricated from a tube radially expanded by at least about400%. 11-20. (canceled)